The present application relates generally to the field of “active implantable medical devices” as defined by Directive 90/385/EEC of 20 Jun. 1990 the Council of the European Communities. The present application more specifically relates to a detection/stimulation microlead intended to be implanted in venous, arterial or lymphatic networks. Such a lead can be used in cardiology (e.g., to be implanted in the coronary sinus vein to stimulate a left or right cavity of the heart). Further, microleads are often useful in many other medical applications, including applications where there is the presence of a venous, arterial or even lymphatic network, including the venous or arterial cerebral network.
Electrical stimulation has led to major advances in neurology in the field of neuromodulation (e.g., a technique to stimulate target areas of the brain for the treatment of disorders such as Parkinson's disease, epilepsy and other neurological diseases). Such a technique often allows a less invasive approach of these treatments and especially superior efficacy of treatments.
It is challenging and difficult to provide stimulation microleads of very small diameter but which are nevertheless extremely robust to provide long-term biostability.
While the size of current implantable leads is typically on the order of 4 to 6 French (1.33 to 2 mm), it would be desirable to reduce the diameter to less than 2 French (0.66 mm). Such a size of microlead could access small veinlets, inaccessible today with larger devices. Such a microlead must also be able to easily navigate through the venous arterial or lymphatic networks with sufficient flexibility to be introduced into vessel networks with high tortuosity, anastomosis, etc.
However, in the prior art, the reduction in lead diameter often increases technological complexity and imposes technical constraints generating risks.
Conventional microleads can include a central conductor for connection to the generator of the implant. The conductor can be coated with an electrically insulating sheath. Such microleads typically include an active portion including one or more detection/stimulation electrodes electrically connected to the central conductor and intended to come into contact with the target vessel wall.
In the prior art, a first production technique of the electrodes for a microlead includes stripping the insulating sheath so as to expose the microcable at one or more points. The stripped points together constitute a network of electrodes connected in series (monopolar lead configuration), allowing multiple points of stimulation and thus providing multizone dissemination of the stimulation energy delivered by the implant. Such a technique is described, for example, in the EP2455131 and US2012/0130464. These applications also, in an alternative embodiment, disclose production of the active part of the microlead by successively and alternately threading on the microcable insulating tubes and short conductive electrodes of platinum-iridium, in the form of rings. The insulating tubes, made, for example, of polyurethane, are affixed to the microcable and the platinum-iridium electrodes are crimped directly to this microcable.
Another technique for forming electrodes on a microcable can include applying a coating to the microcable polyurethane adhesive, leaving locally appearing conductive uncoated surfaces. With techniques that include the exposure of microcable (removal of the insulation or surfaces left in reserve), it is sometimes needed to provide a conductive coating (e.g., an alloy of titanium nitride or a carbon deposit, such as Carbofilm) by a sputtering technique such as described in, e.g., US5370684 and US5387247, to protect the exposed cable to corrosion. Such a coating can be made of an alloy such as MP35N (35% Ni, 35% Co, 20% Cr and 10% Mo) (e.g., considered stainless under standard conditions). However, in certain circumstances, such a material can be relatively sensitive to electrocorrosion—e.g., a corrosion phenomenon accentuated by the current flow in the polar regions (electrodes) and by contact with surrounding body fluids (blood, etc.).
Applicants have identified a need for reducing the risk of infusion of corporeal fluids to the microcable. The microcable can be completely isolated from any contact with the environment of the microlead, hence the additional conductive coating of titanium nitride NiTi or carbon on the electrode areas or the crimping on the microcable of platinum-iridium rings (noble material, resistant to corrosion). This constraint of isolation of the microcable with the external environment, particularly in the region of the electrodes should ideally be respected i) throughout the expected life of the microlead, ten years, and ii) throughout all the mechanical movements of the microlead (e.g., 400,000,000 bending stresses without breaking—corresponding to the average number of heart beats on the life of the microlead). The fulfillment of these conditions is desirable for a microlead intended to be implanted in the body (e.g., in a coronary vessel). A stimulation microlead in the venous system may experience localized of deformation under curvatures much higher than those experienced by a conventional lead, since it must follow the deformation of the veins. This can cause higher microlead stresses than some other application.
U.S. 2009/134134 A1 discloses a lead structure comprising a ring electrode (i.e., band electrode) connected to a coiled conductor embedded in the thickness of the hollow sheath of the lead body. The insulating sheath is locally stripped to discover the conductor, and a ring is welded to the latter for the electric connection. However, in such a lead, the risk of long-term infusion of body fluids toward the conductor remains because no specific mechanism of protection of the joints or isolation of the conductor with the environment of the microlead is provided. U.S. 2006/265037, U.S. 2005/113896, and U.S. Pat. No. 6,018,684 each describe other structures of ring electrodes, having the same drawback.